Position/weight-activated knee locking mechanism

ABSTRACT

A passive knee locking mechanism using the position of the shank and the weight of the user to lock. The mechanism has translational motion, which allows the shank to move linearly along the sagittal/vertical plane, which locks and unlocks the knee, and rotational motion for the shank to swing when unlocked. The knee block, which is disposed at the top of the shank, has a spur gear that locks with a matching gear rack on the top plate of the knee housing. The side plates for the knee housing have a linear path that makes for the translational motion. During heel strike, the stopper in the front plate of the knee housing positions the shank in a straight position. When the user applies their weight for initiating stance phase, the knee block traverses up the path and the spur gear meshes with the matching rack, locking the knee.

CROSS-REFERENCE TO RELATED APPLICATIONS

This nonprovisional application claims priority to U.S. Non-provisionalpatent application Ser. No. 15/131,571, entitled“Position/Weight-Activated Knee Locking Mechanism,” filed on Apr. 18,2016 by the same inventors, which claims priority to U.S. ProvisionalPatent Application No. 62/149,163, entitled “Position/Weight ActivatedKnee Locking Mechanism,” filed on Apr. 17, 2015 by the same inventors,the entirety of each of which is incorporated herein by reference.

FEDERALLY SPONSORED RESEARCH OR DEVELOPMENT

This invention was made with Government support under Grant No.MRI-1229561 awarded by the National Science Foundation. The governmenthas certain rights in the invention.

BACKGROUND 1. Field of the Invention

This invention relates, generally, to transfemoral amputation. Morespecifically, it relates to correction of gait irregularities found intransfemoral prosthesis users.

2. Brief Description of the Prior Art

The human knee is a complex and robust system. It is the most importantjoint for human gait because of its immense load bearing ability. Theloss of such an important joint often makes it difficult for a person toambulate. Because of this and the resulting unnatural application offorces, many transfemoral amputees develop an asymmetric gait that leadsto future complications. Prosthetic knees are required to bewell-designed to cope with all variabilities.

Human gait can be defined as a synchronized and periodic advancement ofeach leg propelling a person forward [100]. It is a complex processinvolving the coordination of various muscle groups belonging todifferent parts of the lower extremity. The balance that holds thecomplex process of human gait together is diminished when a person has alimb amputated. Every part of the lower extremity contributes towards astable gait, especially the joints. The ankle and knee joint areresponsible for load bearing, articulation, and the overall dynamics ofgait [101]. Hence, removing the knee and ankle joints duringtransfemoral amputation severely affects the person's gait [34]. One wayto counteract the changed gait pattern is to improve the prostheticdesign, specifically at the knee joints. Since there are an estimatedseven million transfemoral amputees across the world [86], it isimportant to keep the economics in mind during the design phase so theprosthetic can be low cost and simple, both of which are met by passiveknee mechanisms.

Specifically, the two major phases of human gait are the stance andswing phase [67] [92]; these phases are depicted in FIG. 1. The stancephase is the time the foot is in contact with the ground, whereas theswing phase consists of the time when the foot is swinging in the air.When one leg is going through the stance phase the other leg goesthrough the swing phase. During walking, there also exists a phasecalled double support where both the feet are in contact with theground, which accounts for approximately 10% of the gait cycle. When theheel strike occurs on the stance leg, the swing leg is toeing off duringits pre-swing phase. During the loading response phase, the weightbegins to shift to the stance leg meanwhile the swing leg initiatesswing. During the mid-stance phase, all the load is shifted to thestance leg and the swing leg is in mid-swing. During the terminal stancephase, the stance leg prepares for toe off and the swing leg initiatesheel strike.

An individual's gait is substantially altered when he/she undergoestransfemoral amputation, typically as a result of trauma, accidents, ordue to disease, like diabetes, vascular disease etc. Transfemoralamputation and knee dis-articulation [4] are procedures where the personloses the function of the knee and the ankle joints. With this type ofamputation, a person loses two of the most versatile joints in the humanbody—the knee and ankle. The knee joint is important to human gaitbecause it serves as a junction for the thigh and shank muscles. Theknee locks and unlocks during heel strike and toe off respectively. Kneelocking can be caused either by contraction of muscles (voluntary) or aslight overextension of the knee (involuntary). Without locking of theknee, human legs would buckle and walking would not be possible. Theankle joint is important to human gait because it offers stiffness toavoid collapse of the leg at dorsiflexion or heel strike; at plantarflexion or toe off, it provides control and power to propel the bodyforward. The loss of function of these muscles results in variation ofgait, usually as age progresses [45].

Transfemoral amputees develop a physical asymmetry because of theiramputation, which includes reduced force generation at the knee andankle, reduced control of the leg, and different mass propertiesrelative to their intact leg. Center of pressure of the prosthesis failsto shift towards the posterior during gait initiation [93] and anteriorduring gait termination [94], whereas ideally it should. The physicalchange in the prosthetic leg leads to gait asymmetries that includespatial, temporal, or force differences. Transfemoral amputees with asymmetric prosthesis compensate for altered forces with their intactleg, alter their gait in order to walk comfortably, expend higher energyin walking than able bodied and transtibial amputees, and oftenexperience pain due to compensating forces and torques in the intact legand at the hip joints. There is a correlation between the energyexpenditure and number of joints lost. This functional loss is becauseof the missing joints, or degrees of freedom, in the amputated leg. Toequalize the functional losses, the body has to work harder; in thiscase it is the intact leg that experiences an increase in joint forcemoments and has to expend higher energy [63]. There are also residualstresses that are experienced in the stump as well, resulting indiscomfort while walking [77]. The stresses in the residual stump are abi-product of asymmetric reaction forces and moments. Thus, there is aneed to further refine prosthetic devices.

The foregoing comments regarding transfemoral amputations, transfemoralamputees, and the problems associated therewith and also with theprosthetics, can be applicable to any individual who no longer has fullcontrol of his/her leg from above the knee downwards. This may include,for example, stroke victims as well.

Certain prosthetic devices (including simple and complex prosthetic kneedesigns) and rehabilitation methods are known in the art and attempt tofacilitate these individuals' movement. Stability and control of aprosthesis can be explained with respect to the thigh-knee-ankle (TKA)weight line. As seen in FIG. 2, stability is high and voluntary controlis low when the TKA weight line is anterior to the knee axis;conversely, stability is high and voluntary control is high andstability is low when the TKA weight line is anterior to the knee axis.

For example, U.S. Pat. No. 8,192,501 to Kapelke discusses a prostheticdevice with a knee-lock mechanism, controlling the motion of thearticulating joint. However, the locking mechanism relies on anelectrical signal from a gravity activated tilt-switch. U.S. PatentApplication Publication No. 2010/0292807 to Velez et al. teaches aweight-activated prosthetic knee joint but does not teach any beneficialfoot design. U.S. Pat. No. 6,106,560 to Boender discusses aweight-activated prosthetic knee joint that can be attached to aprosthetic foot. However, it uses hydraulic valves and chambers in itsprosthesis design.

Further, Andrysek et al., Mobility Function of a Prosthetic Knee Jointwith An Automatic Stance Phase Lock, Prosthetic and OrthoticsInternational, 35.2 (2011): 163-170 presents a study developing asimplified automatic stance phase lock (SASPL) mechanism, based on alocking mechanism that engages or disengages depending on loading of theprosthetic limb during weight bearing and provides a securely lockedknee in early to mid-stance phase without restricting knee flexion inpre-swing and swing phase. Shamaei et al., A Quasi-Passive CompliantStance Control Knee-Ankle-Foot Orthosis, IEEE International Conferenceon Rehabilitation Robotics (2013) discusses an orthosis that stabilizesthe knee by implementing a sized spring in parallel with it during theweight acceptance phase of the gait and follows for free (low-stiffness)rotation during the rest of the gait. Trifonov et al., Design Issue andApplications for a Passive-Dynamic Walker, International Journal ofMultimedia and Ubiquitous Engineering, 4.3 (2009): 57-72 teaches adesign for a walker with a knee-locking mechanism. However, each of theforegoing references—and the devices and methodologies discussedtherein—have their limitations and can pose problems for users'stability and control, along with gait symmetry.

Prosthetic knees can be broadly characterized as passive and activemechanisms [102, 10]. Active mechanisms are state of the art and aredesigned to mimic the knee and ankle joint effectively [85]. In manycomparison studies related to walking, such as stair ascent, walking ona slope, and performing ambulatory movements [34, 35, 103], active kneeshave shown lower metabolic strain than passive knees. Many active kneeshave variable settings that allow the user to adjust their prosthetic tothe terrain and condition of their environment. However, all theseadvantages of active knees are expensive and many transfemoral amputeeshave to resort to inexpensive passive knees [104].

There are five kinds of passive knee locking mechanisms, namely: manual,poly-centric, single axis, weight activated, and knee with exteriorhinges [105, 106]. Manual locking mechanisms are generally used byamputees who have minimal capacity for movement, K0-K2 in the amputee Klevels (K is an arbitrary letter assigned by HCFA) [106, 82]. Amputee Klevels are specified to categorize amputees on their ability torehabilitate and is also taken into consideration when choosing aprosthesis. Manual locking allows the amputee to achieve more stabilityfrom the knee joint, since they cannot control the prosthesis in anyother form due to the lack of ambulatory muscles. Poly-centric knees area popular choice for passive knee mechanisms [61]. Poly-centric kneesare generally made of 4, 5, and 6 bar mechanisms [105, 107] where theinstantaneous center of the mechanism shifts during the gait cycle andlocks based on the position of the shank with respect to the thigh inthe gait cycle. Poly-centric knees also offer better control of theswing to the amputee. Single axis systems are simple mechanisms, but arenot as commonly used as poly-centric knees. Weight activated kneemechanisms are often coupled with single axis knees to provide betterlocking [106]. This mechanism utilizes the user's weight to lock theknee during stance phase.

The weight-actuated mechanisms often rely on links that are connectedwith an intricate pattern to either guide high friction surfaces to meshor apply brakes when the weight is acted upon the system. The constantcontact of the components results in high friction leading to more wearof the internal components. Knees with an exterior hinge type mechanismwere used earlier in the development of prosthetic knees and theyresembled an orthotic device.

Accordingly, what is needed is an apparatus that improves the quality ofgait in transfemoral prosthesis users by shifting knee location, thusdecreasing overall prosthesis weight. However, in view of the artconsidered as a whole at the time the present invention was made, it wasnot obvious to those of ordinary skill in the field of this inventionhow the shortcomings of the prior art could be overcome.

All referenced publications are incorporated herein by reference intheir entirety. Furthermore, where a definition or use of a term in areference, which is incorporated by reference herein, is inconsistent orcontrary to the definition of that term provided herein, the definitionof that term provided herein applies and the definition of that term inthe reference does not apply.

While certain aspects of conventional technologies have been discussedto facilitate disclosure of the invention, Applicants in no way disclaimthese technical aspects, and it is contemplated that the claimedinvention may encompass one or more of the conventional technicalaspects discussed herein.

The present invention may address one or more of the problems anddeficiencies of the prior art discussed above. However, it iscontemplated that the invention may prove useful in addressing otherproblems and deficiencies in a number of technical areas. Therefore, theclaimed invention should not necessarily be construed as limited toaddressing only the particular problems or deficiencies discussedherein.

In this specification, where a document, act or item of knowledge isreferred to or discussed, this reference or discussion is not anadmission that the document, act or item of knowledge or any combinationthereof was at the priority date, publicly available, known to thepublic, part of common general knowledge, or otherwise constitutes priorart under the applicable statutory provisions; or is known to berelevant to an attempt to solve any problem with which thisspecification is concerned.

SUMMARY

The long-standing but heretofore unfulfilled need for an improvedprosthetic for individuals suffering from impaired movement in their legfrom above the knee to the foot (due to amputation, stroke, etc.) is nowmet by a new, useful, and nonobvious invention.

In an embodiment, the current invention is a passive prosthetic, such asa passive, asymmetric unilateral transfemoral prosthesis. The passiveprosthetic includes a prosthetic femoral component, a prosthetic shankcomponent, a prosthetic foot component, and a prosthetic knee component.The femoral component is a substantially vertically-oriented shafthaving a top side and a bottom side, where the top side is coupled to aresidual/impaired limb connector (e.g., a knee brace) in underlyingrelation to that connector. The shank component is passive and also hasa top side and a bottom side, where it is rotatable relative to thefemoral component. The foot component of the passive prosthetic is alsopassive and is coupled to the bottom of the shank component inunderlying relation to the shank component. Optionally, the footcomponent can have a passive rollover shape including a radius ofcurvature of a bottom surface of the foot component decreasing toward afront side of the foot component.

The knee component is disposed between the femoral component and theshank component, where the knee component is coupled to the bottom ofthe femoral component in underlying relation to the femoral componentand to the top of the shank component in overlying relation to the shankcomponent. The knee component includes a housing, a spur gear, and agear rack, where the housing substantially encloses the spur gear andthe gear rack within the housing's interior. The spur gear and gear rackmesh with each other when in contact. The spur gear may be a half gearwith one side including teeth and an opposite side being planar;alternatively, the spur gear may be a pair of half gears that correspondto a pair of gear racks.

The knee component has a first locked position and a second unlockedposition. The first position includes the spur gear and gear rack beingmeshed together as a result of the shank component being substantiallyvertical and a user of the device exerting a downward force on the kneeand shank components. In the first position, the knee and shankcomponents are locked in the vertical position. The second positionincludes the spur gear and gear rack having a spaced distancetherebetween (i.e., not meshed) as a result of the user not exerting adownward force on the knee and shank components. In the second position,the shank component can rotate relative to femoral component as in thegait/walking motion of the user. When transitioning between thesepositions, the spur gear and/or gear rack is vertically displaced, forexample by less than about 20 mm.

The shank component and/or the femoral component may have an adjustablelength, for example by including at least two (2) shafts telescopicallyreceived within one another. There may also be included separateextender or coupler shafts to extend the length of the components, ifneeded.

Optionally, the shank component may be coupled to the bottom side of thespur gear, where the spur gear and shank component rotate together whenunlocked. Further, the knee component can include a shaft and bearingsassembly in communication with the spur gear to facilitate therotational motion of the spur gear and shaft component. Additionally, inthis embodiment, the spur gear can be upward/superior facing, and thegear rack can be downward/inferior facing.

In certain embodiments, the apparatus may include one or more stoppersdisposed along a front side of the knee component or the shankcomponent, or both, to prevent the shank component from rotating furtherforward than the vertical locked position, thus preventing an unnaturalbend at the knee. In other words, the stopper(s) prevent any upwardtranslation of the shank component. More specifically, the kneecomponent housing can include a stopper, and the shank component caninclude a stopper, where in the vertical position, the two stopperscontact each other to prevent any further horizontal or verticaldisplacement in an undesired/unnatural direction. These stoppers alsoallow the knee and shank components to assume a precise lockingposition.

In a separate embodiment, the current invention is a position- andweight-activated knee locking apparatus, including the prosthetic kneeapparatus, substantially as described previously.

These and other important objects, advantages, and features of theinvention will become clear as this disclosure proceeds.

The invention accordingly comprises the features of construction,combination of elements, and arrangement of parts that will beexemplified in the disclosure set forth hereinafter and the scope of theinvention will be indicated in the claims.

BRIEF DESCRIPTION OF THE DRAWINGS

This patent or application file contains at least one drawing executedin color. Copies of this patent or patent application publication withcolor drawings will be provided by the Office upon request and paymentof the necessary fee.

For a fuller understanding of the invention, reference should be made tothe following detailed description, taken in connection with theaccompanying drawings, in which:

FIG. 1 depicts the eight (8) phases of human gait. The stars indicateheel strike at initial contact and knee strike at terminal swing.

FIG. 2A shows a short stump with high stability and low voluntarycontrol. The thigh-knee-ankle (TKA) weight line is in the anterior tothe knee joint.

FIG. 2B shows a medium stump with medium stability and medium voluntarycontrol. The TKA weight line is in the middle to the knee joint.

FIG. 2C shows a long stump (knee disarticulation) with low stability andhigh voluntary control. The TKA weight line is in the posterior to theknee joint.

FIG. 3 is a prosthetic simulator according to an embodiment of thecurrent invention.

FIG. 4 depicts human knee locking and unlocking during walking. Thefigure shows how the tibia and patella rotate as well as translate overthe femur.

FIG. 5A depicts a prosthetic knee mechanism with two degrees of freedom,vertical translation and single axis rotation, where the knee mechanismis shown in an unlocked position.

FIG. 5B depicts the prosthetic knee mechanism of FIG. 5A, where the kneemechanism is shown in a locked position.

FIG. 6A depicts a mechanism of the knee, specifically the toe off phaseduring knee unlocking.

FIG. 6B depicts a mechanism of the knee, specifically the initial swingphase during knee unlocking.

FIG. 6C depicts a mechanism of the knee, specifically swing during kneeunlocking.

FIG. 6D depicts a mechanism of the knee, specifically swing during kneeunlocking.

FIG. 6E depicts a mechanism of the knee, specifically swing during kneelocking.

FIG. 6F depicts a mechanism of the knee, specifically swing during kneelocking.

FIG. 6G depicts a mechanism of the knee, specifically the stopper strikephase during knee locking.

FIG. 6H depicts a mechanism of the knee, specifically the terminal swingphase during knee locking.

FIG. 7 depicts gait with asymmetric prosthetic simulator, as illustratedwithin a simulation.

FIG. 8A depicts a prosthetic thigh according to an embodiment of thecurrent invention, where the figure includes exemplary length settings.

FIG. 8B depicts a prosthetic shank according to an embodiment of thecurrent invention, where the figure includes exemplary length settings.

FIG. 9 depicts a prosthetic foot according to an embodiment of thecurrent invention.

FIG. 10 depicts a mechanism of foot rollover during walking.

FIG. 11A depicts a rollover shape of the prosthetic foot, according tocertain embodiments of the current invention, where the design has alarge constant radius anterior to the ankle line and a smaller radius tothe posterior.

FIG. 11B depicts a rollover shape of the prosthetic foot, according tocertain embodiments of the current invention, where the design has aconstant decreasing radius towards the anterior and a constantincreasing radius to the posterior.

FIG. 12A is a foot shape polar plot.

FIG. 12B is an applied and reaction force plot.

FIG. 13A is a graphical illustration of a length of a normal gait step.

FIG. 13B is a graphical illustration of a swing time of a normal gaitstep.

FIG. 13C is a graphical illustration of a length of a low knee settinggait step.

FIG. 13D is a graphical illustration of a swing time of a low kneesetting gait step.

FIG. 13E is a graphical illustration of a length of a medium kneesetting gait step.

FIG. 13F is a graphical illustration of a swing time of a medium kneesetting gait step.

FIG. 13G is a graphical illustration of a length of a high knee settinggait step.

FIG. 13H is a graphical illustration of a swing time of a high kneesetting gait step.

FIG. 13I is a graphical illustration of step lengths of a CAREN system.

FIG. 13J is a graphical illustration of swing times of a CAREN system.

FIG. 14 depicts stress analysis of an exemplary prosthetic kneeaccording to an embodiment of the current invention.

FIG. 15A depicts a stress analysis of an exemplary prosthetic thighaccording to an embodiment of the current invention. The figure showsthe prosthetic thigh assembly with the top of the small cylinderconstrained and ground reaction force of 3000N acting on the bottomcircumference of the large cylinder.

FIG. 15B depicts a stress analysis of an exemplary prosthetic shankaccording to an embodiment of the current invention. The figure showsthe prosthetic shank assembly with the upper circumference of the largecylinder fixed and a ground reaction force of 3000N acting on the bottomof the small cylinder.

FIG. 16A depicts a stress analysis of an exemplary prosthetic footdesign according to embodiment of the current invention. The figureshows Von Mises stress experienced by the constant radius design of FIG.11A.

FIG. 16B depicts a stress analysis of an exemplary prosthetic footdesign according to embodiment of the current invention. The figure VonMises stress experienced by the shape constant decreasing radius designof FIG. 11B.

FIG. 17 depicts stress analysis of an exemplary base plate according toan embodiment of the current invention.

FIG. 18A depicts a full body marker layout.

FIG. 18B depicts markers for the prosthetic, according to an embodimentof the current invention.

FIG. 19 depicts gait with prosthetic simulator in a VICON model.

FIG. 20A is a cumulative step length plot for all subjects.

FIG. 20B is a cumulative swing time plot for all subjects.

FIG. 21A is a graphical illustration of step length for the CARENsystem.

FIG. 21B is a graphical illustration of swing time for the CAREN system.

FIG. 22 a ground reaction force graph for every trial on the CARENsystem.

FIG. 23A depicts a modular prosthetic simulator according to anembodiment of the current invention.

FIG. 23B depicts a module prosthetic leg according to an embodiment ofthe current invention.

FIG. 24A depicts a knee assembly without collars, according to anembodiment of the current invention.

FIG. 24B depicts the knee assembly of FIG. 24A in a dismantleddisposition.

FIG. 24C is a knee assembly dismantled isometric view.

FIG. 25A is an aluminum raw material 2D drawing.

FIG. 25B is a spur gear and gear rack 2D drawing.

FIG. 26 depicts a prosthesis, according to an embodiment of the currentinvention, used to test the position and weight-actuated knee lockingmechanism.

FIG. 27 is an exploded view of the prosthesis of FIG. 26

FIG. 28A depicts a knee position where the knee is locked when the shankgear meshes with the femoral gear, which happens because of the user'sweight acting on the knee.

FIG. 28B depicts a knee position where the knee unlocks when the user'sweight does not act on it, and the shank gear disengages from thefemoral gear and slides down the slot.

FIG. 28C depicts a knee position where the shank rotates about thebearing as the user swings their residual limb.

FIG. 28D depicts a knee position where as the user reaches the end oftheir swing, the shank swings back like a pendulum and hits the stopperto assume the position for locking. The locking cycle then begins as theuser applies their weight on the prosthesis. In FIGS. 28A-28D, the redarrows indicate the application of force and green arrows indicatemotion.

FIG. 29A depicts an example of walking with the prosthesis, specificallythe phase with heel strike.

FIG. 29B depicts an example of walking with the prosthesis, specificallythe phase with knee strike/end of swing.

FIG. 29C depicts an example of walking with the prosthesis, specificallythe phase with full swing.

FIG. 29D depicts an example of walking with the prosthesis, specificallythe phase with initial swing.

FIG. 29E depicts an example of walking with the prosthesis, specificallythe phase with the shank unlocking and sliding down.

FIG. 29F depicts an example of walking with the prosthesis, specificallythe phase with toe-off.

FIG. 29G depicts an example of walking with the prosthesis, specificallythe phase with stance/loading response.

FIG. 29H depicts an example of walking with the prosthesis, specificallythe phase with knee locking after heel strike.

FIG. 30A depicts a tracing of the motion of the knees at baseline normalwalking.

FIG. 30B illustrates knee motion with the prosthesis on the right leg.The prosthetic knee was slightly lower than the normal knee toaccommodate the prosthetic simulator.

FIG. 31A depicts knee angles for normal walking recorded as “baseline.”

FIG. 31B depicts knee angles with the prosthesis on the right leg,showing that the shank of the prosthesis has a larger knee flexion angleand the intact knee compensates by keeping the flexion to a minimum.

FIG. 32A depicts angular velocity of the knees during normal walkingrecorded as “baseline.”

FIG. 32B depicts angular velocity of the knees with the prosthesis onthe right leg, showing that the shank of the prosthesis has a largermagnitude of angular velocity from flexion to extension and vice versa,while the intact knee compensates with low angular velocity and a longerstance phase.

DETAILED DESCRIPTION

In the following detailed description, reference is made to theaccompanying drawings, which form a part thereof, and within which areshown by way of illustration specific embodiments by which the inventionmay be practiced. It is to be understood that other embodiments may beutilized and structural changes may be made without departing from thescope of the invention.

As used in this specification and the appended claims, the singularforms “a”, “an”, and “the” include plural referents unless the contentclearly dictates otherwise. As used in this specification and theappended claims, the term “or” is generally employed in its senseincluding “and/or” unless the context clearly dictates otherwise.

It should be understood that any reference to an element herein using adesignation such as “first,” “second,” and so forth does not limit thequantity or order of those elements, unless such limitation isexplicitly stated. Rather, these designations may be used herein as aconvenient method of distinguishing between two or more elements orinstances of an element. Thus, a reference to first and second elementsdoes not mean that only two elements may be employed there or that thefirst element must precede the second element in some manner. Also,unless stated otherwise a set of elements may comprise one or moreelements.

The phrases “connected to” and “coupled to” refer to any form ofinteraction between two or more entities, including mechanical,electrical, magnetic, electromagnetic, fluid, and thermal interaction.Two components may be connected or coupled to each other even thoughthey are not in direct contact with each other. For example, twocomponents may be coupled to each other through an intermediatecomponent.

The directional terms “proximal” and “distal” are used herein to referto opposite locations on a medical device. The proximal end of thedevice is defined as the end of the device closest to the practitionerwhen the device is in use by the practitioner. The distal end is the endopposite the proximal end, along the longitudinal direction of thedevice, or the end furthest from the practitioner.

In an embodiment, the current invention is a passive knee mechanism thatincorporates both linear motion and rotary motion of the prostheticshank. This knee design closely mimics the human knee kinematics. Thesystem is inexpensive and has the potential to be precursor tobiologically-inspired transfemoral prosthetic knee designs, both passiveand active. A passive mechanism was tested successfully on asymmetrictransfemoral prosthesis.

In an embodiment, the current invention is a passive knee lockingmechanism that relies on the position of the shank and the weight of theuser to lock. The mechanism has translational motion, which allows theshank to move linearly along the sagittal plane (up and down), whichlocks and unlocks the knee, and rotational motion for the shank to swingwhen unlocked. The knee block, which is at the top of the shank has ahalf gear that locks with a matching gear rack on the top plate of theknee housing. The knee block also has a slot for a shaft and issupported by bearings on the side plates of the knee housing. The shaftand bearings combined allow the rotational motion. The side plates forthe knee housing have a linear path that makes for the translationalmotion. During heel strike, the stopper in the front plate of the kneehousing positions the shank in a straight position, and when the userapplies their weight for initiating stance phase, the knee blocktraverses up the path and the half gear meshes with the matching rackthat locks the knee. The path through which the shank traverses can bemodified to suitably mimic the kinematics of the human knee.

In an embodiment, the current invention is a passive prosthetic deviceand system that generally includes a passive knee, a prosthetic thighand shank, a passive foot, and a knee brace, as seen in FIG. 3. Thedevice improves both energy costs and gait symmetry for transfemoralamputees, which can also be understood to encompass individuals who haveimpaired movement in their lower extremities. The prosthesis hasadjustable knee locations for each wearer and can be lighter than ahuman shank. As the knee location shifts downwards, the moment arm ofthe shank decreases, therefore, having a shorter shank swing phase. Thereduction in weight combined with a shorter shank swing reduces theenergy cost required to walk [41].

Example 1

Prosthetic Knee

The depiction in FIG. 4 shows that the tibia and patella performtranslational motion, as well as rotational motion, over the femoralsurface. The cartilage acts as a guiding path for the patella. Thelocking and unlocking mechanism of the knee during normal walking iscompletely passive, utilizing only the dynamic forces of forward motion.The knee can be locked in other positions apart from the extensionposition by activating muscles to lock the knee in place. However,because certain embodiments of the current invention are completelypassive, the invention would incorporate locking only when the knee isin the fully extended position, where the user/operator's weight pushesthe system down to lock. The knee mechanism experimented within thisprosthesis design provides for a position/weight-activated lockingmechanism, as will become clearer as this specification continues.

The knee was designed with the intent to incorporate translational androtational motion in the mechanism. The position-/weight-activatedlocking mechanism simplifies the complex trajectory by allowing only twodegrees of freedom, vertical translation, and single axis rotation tothe complete system, depicted in FIGS. 5A-5B. The vertical translationis achieved by a simple vertical slot allowing the knee joint to have asmall vertical displacement (e.g., about 15 mm or up to about 20 mmuntil the change in gait becomes excessively noticeable or uncomfortablefor the user), as can be seen in in FIG. 5A. This translational freedomis utilized to lock the knee (FIG. 5B) and unlock the knee (FIG. 5A).

The knee is locked when the weight of the wearer acts upon the kneeassembly making the knee assembly reach its upper limit of its verticaltranslation (see FIGS. 6A and 6H). The locking is carried out by thehalf spur gear meshing with the gear rack fitted on top of the kneehousing, depicted in FIGS. 6A and 6H. The knee starts to unlock as thewearer releases their weight at toe off (FIGS. 6A-6C), and the force ofgravity pulls the knee joint to the lower limit of its verticaltranslation, depicted in FIG. 6D. The single axis rotational freedom isprovided to help the knee and shank perform a natural swing. The knee isfree to rotate as its vertical position changes downward and upward, solong as the half spur gear is not meshed with the gear rack (see FIGS.6C-6F). Just before the terminal swing phase, the knee strikes thestopper to attain its position before the upward translation, as seen inFIG. 6G. At this point, the user's weight locks the knee by meshingtogether the half spur gear and gear rack, as seen in FIG. 6H. Aschematic with the gait with an asymmetric prosthesis is shown in FIG.7.

Regarding the specifications of the prosthetic knee, generally a knee isrequired to overcome shock and transient loads generated during humanwalking. The maximum load on the prosthetic simulator can be estimatedto be as much as three (3) times the wearer's bodyweight [50]. Assumingthe wearer's weight to be a maximum of 100 Kg, the dynamic transientforces, which are higher than the heel strike force, is about three (3)times the person's bodyweight. Therefore, every component of the kneeshould be able to withstand close to 3000N of force at any given point.The knee is designed to be effectively heavier than the prosthetic thighand shank, as seen in Table 1. This ensures that the center of mass ofthe prosthesis is near the knee. This allows the center of mass tochange based on the position of the knee in the prosthesis. The changein center of mass is an important observation provided by Sushko et al.[88], which explains that a symmetric gait can be achieved by varyingthe center of mass of the knee to asymmetric locations with counterweights on the intact leg.

TABLE 1 Mass of prosthetic components. Combined S. Weight Number ofWeight No Pad (Grams) Components (Grams) 1 Lower Small Cylinder 207.9 1207.9 2 Upper Small Cylinder 223.8 1 223.8 3 Front Knee Plate + Stopper129.1 1 129.1 4 Steel Gear Rack 38 2 76 5 Upper Large Cylinder 269.7 1269.7 6 Knee Top Plate 90.3 1 90.3 7 Lower Large Cylinder 255.4 1 255.48 Knee Side Plate 80.6 2 161.2 9 Ball Bearing 60.5 2 121 10 Steel Shaft241.4 1 241.4 11 Collar 89 2 178 12 Half Gear 87.4 2 174.8 13 Knee Block276.5 1 276.5 14 Brass Connecting Bolt 33.1 1 33.1 15 Aluminum Bars 1464 584 16 Line Holder 62.4 4 249.6 17 Right Angle Bracket 118.2 2 472.818 Base Plate 281.3 1 281.3 19 Bolt and Nut 26.3 22 578.6 Weight of KneeBrace: 2.19 Kg Weight of Extenders: 0.95 Kg Weight of Knee: 1.5 KgWeight of Total system: 4.6 Kg

Though any suitable material is contemplated herein, certain embodimentsof the prosthetic knee may be formed of a combination of aluminum andsteel. Aluminum can be used to build the knee housing, knee block, andcollars. Aluminum is readily available, easy to machine, and has a highstrength-to-weight ratio. However, steel has a higher shear strength,and hence steel can be used to build the gear, gear rack, and shaft. Thegear, gear rack, and shaft each experience shear and shock loads everygait cycle, and aluminum might fatigue easily under such conditions. Assuch, the gear, gear rack, and shaft can be formed of a strongermaterial than the knee housing, knee block, and collars.

Prosthetic Thigh and Shank

The thigh and shank are adjustable linkages. They are used to adjust theheight of the prosthesis to allow different knee locations. They arecylinders formed of any suitable material (e.g., aluminum), where thereis one hollow cylinder that adjustably and telescopically receives onesolid cylinder. The links are fit in place with the help of a bolt orother suitable mechanism to stably adjust total length.

The thigh and shank are similar in design, each including two cylinders.The thigh and shank are each adjustable to an array of settings; forexample, each can have six settings, thus permitting a total of twelvesettings for the length of the prosthesis. Each setting can give theuser a predetermined difference in length (e.g., about 20 mm), depictedin FIGS. 8A-8B. FIG. 8A depicts the difference in lengths when the kneeis locked, and FIG. 8B depicts the difference in length when the knee isunlocked. As can be seen, knee locking and unlocking changes length byabout 15 mm (or otherwise equal to the vertical displacement of the spurgear within the knee mechanism).

In an embodiment, there is a constant length of the knee (Knee TopPlate+Gear Rack Width+Knee Block) and the foot height equivalent toabout 105 mm or other predetermined height. Therefore, depending on theanatomical thigh and shank length of the wearer, the settings can beadjusted, keeping in mind the constant lengths of the knee and foot.

Images of an embodiment of the knee locking mechanism is presented inFIGS. 24A-24C. It can be seen in FIG. 24A that there may be two (2) spurgears (see also FIG. 24B) corresponding to two (2) gear racks (see alsoFIG. 24C). Alternatively, there can be a single spur gear thatcorrespond to a single gear rack (see knee mechanism of FIG. 27, whichwill be described in more detail as this specification continues). FIG.25A is a schematic of the leg and knee components, and FIG. 25B is aclose-up of the spur gear and gear rack interaction, where they are notmeshed in the figure, but it can be seen how they would mesh/locktogether when the spur gear and/or gear rack traverse/displace acrossthe slot.

Foot

The foot designs of passive dynamic walkers offered an insight intopassive foot designs that can be employed in the current prostheticsimulator. Passive dynamic walkers have been modeled with a point foot,curved foot, and in more advanced biped walkers, ankles that can provideforces similar to dorsiflexion. The point foot is an easy analyticalmodel that was used by Chen [8] to analyze the five mass model. Thecurved foot model was first proposed by McGeer [56], that is a constantradius foot with a radius approximately one third leg length. This footshape allowed the PDW's legs to clear the ground easily. The foot designby Honeycutt explored the possibilities of testing changing radius footdesigns in passive dynamic walkers, which would enable the foot torelease the energy stored during heel strike at toe off [40]. In anotherstudy, they showed that constant radius foot designs can be replacedwith flat foot designs that were mounted on the ankles using torsionalsprings [95].

Contrastingly, in certain embodiments of the current invention, thedesign of the foot mechanism was maintained as simple as possible,resulting in a foot that does not require an ankle, though an ankle iscontemplated by the current invention. The foot design was based onrollover shapes; an example of the rollover-shaped foot can be seen inFIG. 9. Rollover shapes are defined by the change in center of pressureof the foot during walking (see FIG. 10). There are three main phaseswhen the center of pressure at the foot changes: stride initiation,steady state, and termination. In this design, a constant radiusrollover shape, which is one-third of the total leg length, isconsidered [58]. The assumption is made based on the rollover shapesanalyzed in [31, 58, 67], which show that most of the rollover happensin the anterior of the foot.

Two (2) rollover shapes, depicted in FIGS. 11A-11B, were tested for thefoot rollover shape used in the current prosthesis, though alternativesuitable foot designs are contemplated herein as well. The rollovershape for the first design (FIG. 11A) has a constant radius which isone-third of the leg length. This constant radius abruptly changes to asmaller constant radius near the posterior of the foot. This design didnot necessarily succeed because it tends to roll backwards towards thesmaller radius, which made it difficult to walk forward.

The second foot design (FIG. 11B) is based on the kinetic shape concept[26]. Kinetic shapes roll on flat surfaces when a force is applied onits axle. This foot shape will roll forward when a person applies theirweight on it. This emulates dorsi-flexion of an ankle joint; the shapecompensates for the lack of an ankle joint. The shape for this specificfoot shape is also shown in FIG. 9. Equation 2 is the vertical force,which is based on the assumption that a person weighing 100 kg will usethe foot. Equation 3 is the horizontal force that will be generated whenthe force is applied. Equation 4 is the initial radius assumed for theshape based on one third leg length. These variables are substitutedinto Equation 1, resulting in Equation 5. The foot design that wasultimately implemented in the study is shown in FIG. 11B, where theradius decreases towards the front of the foot. This design workedsuccessfully and was implemented in the final design (see FIG. 9) usedfor testing.

$\begin{matrix}{{R(\theta)} = {\exp\left\lbrack {{\int{\frac{F_{r}(\theta)}{F_{v}(\theta)}\ d\;\theta}} + {Constant}} \right\rbrack}} & {{Eq}.\mspace{14mu} 1} \\{{F_{v}(\theta)} = {1000\mspace{14mu} N}} & {{Eq}.\mspace{14mu} 2} \\{{F_{r}(\theta)} = {50\mspace{14mu} N}} & {{Eq}.\mspace{14mu} 3} \\{{{Initial}\mspace{14mu}{radius}} = {0.30\mspace{14mu} m}} & {{Eq}.\mspace{20mu} 4} \\{{R(\theta)} = {{\exp\left\lbrack {\int{\frac{50\mspace{14mu} N}{1000\mspace{14mu} N}d\;\theta}} \right\rbrack}_{\theta = 0}^{\theta = \pi} + {0.30\mspace{14mu} m}}} & {{Eq}.\mspace{14mu} 5}\end{matrix}$

The foot assembly is fit rigidly to the lower solid cylinder of theprosthetic shank. The foot does not need an ankle mechanism because ofkinetic shape rolls forward when the user applies their force on it. Asthe user's moment of inertia shifts forward, the foot starts to rollinto a smooth forward motion leading to toe off (see FIG. 10).

Knee Brace

In an embodiment and for illustration and testing purposes herein, thetransfemoral prosthetic simulator is specifically designed for anon-amputee wearer. The design utilizes an interface between theprosthesis and the wearer's leg. The wearer's leg is held at asubstantially right angle and secured tightly. The prosthesis isdesigned to be fit on the knee brace by locking it on a bolt.

The knee brace should be light weight, rigid, durable, and comfortable.This design may be formed completely of aluminum and the differentcomponents are secured with steel bolts, depicted in FIG. 3. Aluminummay be used because it is easy to machine and has a higher shearstrength than acetal resin, plastic, and wood. The frame is rigid inorder to restrain the movement of the wearer's leg because it caninterfere with the motion of the prosthesis. Aluminum used for the bracealso is light and strong enough to withstand continuous load cycles.Table 2 shows the safety factors for the base plate which is thecomponent that will experience the maximum load.

TABLE 2 Minimum factor of safety for prosthetic components. S. No.Prosthetic Component Minimum Factor of Safety 1 Prosthetic Knee 1.4 2Prosthetic Thigh 1.5 3 Prosthetic Shank 2.6 4 Foot 7.2 5 Base Plate 1.4

The brace also defines the position of the prosthetic leg with respectto the Thigh-Knee-Ankle line, discussed previously. The prosthetic legis placed to the anterior of the wearer's Thigh-Knee-Ankle line toensure high control and low stability of the prosthesis. This can beimportant because the wearer in this case has all the muscles intact intheir leg. The positioning may change depending on the needs of the userfor the prosthesis.

Two unique problems arose while testing the knee brace. First, the kneebrace tends to slip down. The solution to this problem was to secure thebrace to the safety harness using bungee cords or other suitablesecuring apparatus or means as is known in the art. The second problemwas comfort. The metal components were hard and would hurt the wearerwhile walking. To ensure the comfort of the wearer, they wore flexibleknee braces that provide enough padding. In addition to the paddedflexible knee braces, the metal knee brace can be provided withadditional padding for comfort.

It is thus contemplated herein that the device can include simple rigidframe with adjustable shank and thigh lengths to accommodate a widerange of users. The frame may also have an acute angle to have moreclearance of the bent leg. The padding can be designed in such a waythat it can arrest the user's leg movement.

Safety Analysis

All components were subjected to a maximum load of 3000N, assumed forthree times the body weight of a person weighing 100 kg [50]. Althoughthe load may be shared by the components, the tests were to make surethat all components will not fail when maximum load is applied. The loadsimulations were carried out in SOLIDWORKS SIMULATIONXPRESS package. Theprosthetic components were fixed at points according to their designconstraints, and a force of 3000N was applied on the components in thedownward direction, for circular components it was applied on thesurface. The testing was carried out for the aluminum and delrincomponents because they have a higher chance of failure than steel. Thisis because steel has a very high shear strength compared to aluminum anddelrin. Table 2 shows the factor of safety of the components that aresubject to the forces directly. Graphical illustrations of componentanalysis can be seen in FIGS. 13A-13J (prosthesis worn on right leg),comparing knee heights with respect to step lengths and swing time.

Analysis was also conducted on assemblies. The main assemblies that wereconcentrated upon were the knee, foot, thigh, shank, and base plate. Theknee assembly stress analysis depicted in FIG. 14 shows that the topplate of the knee, ends of the shaft, and the locking gear areconstrained, and a ground reaction force of 3000N is applied to the kneeblock and to the meshing portion of the gear rack. This results in highstress in the top plate, shaft and knee block due to high force actingupon them. The factor of safety is low because the knee block is madeout of aluminum and it experiences a high shear force. The system on thewhole, however, proves that the design was correct and would not failunder the given load.

The prosthetic thigh and shank analysis were performed under similarconditions as the knee. FIGS. 15A-15B depict the thigh and shankVon-Mises stresses, respectively. In the prosthetic thigh assembly, thetop of the upper small cylinder is constrained and a force of 3000N isapplied on the bottom of the upper large cylinder. This simulates theforce coming from the top plate of the knee to the upper large cylinder.On the prosthetic shank assembly, the loading is the reverse of thethigh. As seen, the maximum stresses are around the holes. The safetyfactor of the shank is higher because of the force that acts upon thesolid cylinder.

Analysis was performed on both foot designs of FIGS. 11A-11B; the stressanalyses are depicted in FIGS. 16A-16B. The top surface of the foot isconstrained in order to avoid the calculation of maximum displacement,which is not important for this trial. A ground reaction force of 3000Nis applied on both the designs and it is observed that the factor ofsafety of the second design (FIGS. 11B and 16B) is higher than the first(FIGS. 11A and 16A). This change can be explained because of the shapesof the foot designs. The first design (FIGS. 11A and 16A) has twoconstant radii that change the rollover shape suddenly at the ankleline. In the second design (FIGS. 11B and 16B), which has a constantdecreasing radius, the rollover shape is gradual and is capable ofdistributing the force better than the constant radius design.

Stress analysis was also performed on the base plate with the thighassembly attached, as can be seen in FIG. 17. The attachment bolt holeis constrained and a ground reaction force of 3000N is applied on thebottom of the large cylinder. As seen in FIG. 17, the hole experienceshigh stress concentration. The safety factor indicates that the designis still far from failing.

Testing/Study

Previous research has shown that a passive dynamic walker (PDW) with analtered knee location can exhibit a symmetric step length. Theasymmetric prosthesis demonstrated herein aims to find a balance betweenthe different types of asymmetries to provide a gait that is moresymmetric and to make it easier overall for an amputee to walk.

A passive, asymmetric unilateral transfemoral prosthetic simulator wasdeveloped to emulate this PDW with an altered knee location. Theprosthetic simulator designed for this research had adjustable kneesettings simulating different knee locations. The prosthetic simulatorwas tested on able-bodied participants with no gait impairments, thuseliminating the compatibility problems that come with testing it onamputees. The kinetic and kinematic data was obtained using a VICONmotion capture system and force plates.

This research analyzed the kinematic and kinetic data with differentknee locations (high, medium, and low) and normal walking. This data wasanalyzed to find the asymmetries in step length, step time, and groundreaction forces between the different knee settings and normal walking.The current prosthesis differs significantly from the conventional artbecause the knee location in the device is below the anatomicalposition, thus allowing for a design without an offset to accommodatethe knee. The shifting of the knee also makes the design lightweight.

The study showed that there is symmetry in step lengths for all thecases in overground walking. The knee at the lowest setting was theclosest in emulating a normal symmetric step length. The swing times foroverground walking showed that the healthy leg swings at almost the samerate in every trial and that the leg with the prosthetic simulatoreither can be symmetric, like the healthy leg, or can have a higherswing time. Step lengths on the treadmill also showed a similar pattern,and step length of the low knee setting were the closest to the steplength of normal walking. The swing times for treadmills did not show asignificant trend. Kinetic data from the treadmill study showed thatthere was force symmetry between the low setting and normal walkingcases. In conclusion, these results show that a low knee setting in anasymmetric prosthesis may bring about spatial and temporal symmetry inamputee gait. This study was important to demonstrate that asymmetriesin amputee gait can be mitigated using a prosthesis with a knee locationdissimilar to that of the intact leg. Tradeoffs can be made to achievesymmetric step length, swing times, or reaction forces.

The scope of this study is to demonstrate the efficacy of the prosthesiswith the shifted knee location in real world conditions. The prostheticsimulator is designed to have variable knee locations. The walkingbehavior of the wearer is compared at every knee location. The look ofthe prosthesis is out of the ordinary because of the asymmetry. Benefitsof this type of prosthesis is that it is an inexpensive system, canprovide a comfortable gait, and can reduce energy costs of the user.

Kinetic and Kinematic Data Acquisition

The data acquisition was done in two stages. The first stage was tocollect kinematic data on three subjects. The data was collected fortrials consisting of normal walking, knee at high setting, knee atmedium setting, and knee at lowest setting. A VICON system which has anaccuracy of 1 mm and sampling rate of 120 Hz was used to obtain thekinematic data. The VICON uses infrared cameras and reflective markersto accurately track every marker's motion in 3D space. As depicted inFIG. 18A, markers were placed on the lower extremity of the participantsfor normal walking, and as depicted in FIG. 18B, markers were placed onthe lower extremity of the participants with the prosthetic simulator.

The three participants chosen for the study were all male and did nothave any gait disability. All participants wore the prosthesis on theirright leg. The kinematic data was collected on all subjects for all fourtrials. Table 3 shows the height, weight, leg length, and shank length.The participants walked on a wooden platform, which is the same asoverground walking. All participants followed an approved University ofSouth Florida Internal Review Board (IRB) protocol.

TABLE 3 Participant Data Leg Shank Weight Height Length Length Subject(Kilograms) (Centimeters) (Centimeters) (Centimeters) 1 96 186 98 43 285 189 110 55 3 108 184 98 52 Average 96.3 186.3 102 50 Standard 11.52.5 6.9 6.2 Deviation

The second stage was to collect kinetic and kinematic data. The trialsfor this data acquisition are the same as the first stage. A ComputerAssisted Rehabilitation Environment (CAREN) system was used for thisstage of testing. The CAREN system also has a VICON system for kinematicdata acquisition. It also has a split-belt treadmill with force platesto measure the kinetic data. Therefore, the first stage is groundwalking and the second stage involves treadmill walking.

Subject 3 from the first stage was put through the complete set oftrials (normal walking, knee at high setting, knee at medium setting,and knee at lowest setting) for the second stage. The kinetic andkinematic data were recorded for both legs. The kinematic data for bothstages were post processed to obtain useful data. The gait cycle of theskeletal frame from the processed data for the gait is shown FIG. 19.

Kinematic Data Analysis for First Stage Testing

The step lengths were calculated by finding the difference between theposition of the left heel and the right heel. Corresponding swing timeswere also found for each leg. As shown in FIGS. 20A-20B, step length andswing time graphs were plotted for each trial, and a cumulative averageand standard deviation was found for all subjects. Step lengths for alltrials were symmetric. The step lengths of the knee at low setting werehigher than the other settings. The medium setting was very close to thestep lengths of the low setting. The knee at high setting was overallthe farthest from normal walking step length. Therefore, the steplengths of the prosthesis at low setting were symmetric and seen to beclosest to normal walking.

The swing times were also found to be asymmetric. The left leg (healthyleg) had the same time for its swings for all cases. The leg with theprosthetic simulator was seen either to have a swing time symmetric tothe healthy leg or to take more time to swing. Therefore, the cumulativestep times do not show a pattern such as the one observed in steplengths.

Kinematic and Kinetic Data Analysis for Second Stage Testing on CARENSystem

Kinematic data obtained from the CAREN system is shown in FIGS. 21A-21B.Referring to FIG. 21A, it can be seen that the subject took longer stepswith the prosthetic leg; note that the swing time of the prosthesis(FIG. 21B) was shorter than the left leg. Step lengths of the left legwere slightly asymmetric in the medium knee setting and in the high kneesetting. The low knee setting was the closest trial to emulate normalwalking gait. The high knee setting showed the most asymmetry in steplengths. Swing times of the low knee setting was symmetric, whereas thehigh knee setting is asymmetric.

Now referring to FIG. 22, kinetic data of the ground reaction forces wasalso obtained for all the trials on the CAREN system. Looking at thedata, it is apparent that the user applied more force on the prosthesis.This is because there is a high shock load that is applied on theprosthesis, this is another form of compensation mechanism. Aninteresting symmetry in the forces is seen in the high position; thesubject was walking with slightly more force on the healthy leg than onthe prosthesis.

The most significant trend observed is that the closest force thatresembles normal walking was the trial with the low knee setting.

Comparison of Data from Treadmill Walking and Overground Walking

The data acquired from the ground walking tests was small compared tothe data acquired from the treadmill. This was because of the spatialconstraints of a VICON system, which cannot cover a very large distancefor ground walking. This was not an issue with a treadmill, as moresteps can be recorded on a treadmill. Previous research on healthysubjects has not shown any statistically significant differences betweenoverground walking and treadmill walking [76]. A recent study whichcompared overground walking with treadmill walking in the CAREN systemof transtibial amputees and healthy subjects showed that overground andtreadmill walking were similar enough, except for a slight variabilityin step width and step time results [17].

As discussed, the data was analyzed from ground walking and treadmillwalking. The data for step length showed that overground step lengthswere more symmetric compared to the treadmill step lengths. The swingtimes for the left leg varied from ground walking to treadmill, but theprosthesis had almost the same times in both cases.

DISCUSSION

The asymmetric transfemoral prosthetic simulator has its concept rootedin the passive dynamic walker (PDW) model. The asymmetric PDW model inthis case can simulate variation in knee location, thigh width, shankwidth, damping and stiffness of the knee, and leg lengths.

FIGS. 23A-23B depict modular embodiments of the current invention usingthe data and conclusions reached above. These embodiments are fittedwith symmetric locks that can be opened when the person needs to sit.Modularity allows the user to have multiple positions of the kneelocation. Modularity also allows for sufficient room to customize theprosthesis. Another interesting use of the prosthetic simulator can be aform of hands free crutches with knee joints. This will allow the crutchuser to have a more natural gait.

An asymmetric transfemoral prosthesis that can be fitted on amputees isalso contemplated by the current invention, including a neural feedbackcomponent.

Example 2

Described herein is the design and preliminary testing of a novelpassive position and weight activated knee locking mechanism for use inlower limb prosthetics. The mechanism utilizes the dynamics of the userto lock the knee during stance and unlock during the swing phase.Results from testing the knee mechanism show trends that are differentfrom a normal human knee, which is to be expected. The prosthetic kneeis designed to have low friction during swing of the shank and, hence,the flexion and extension angles and angular velocities are largercompared to a normal knee. The kinematics show a cyclic trend that ishighly repeatable.

Structure/Design

The position and weight activated knee locking mechanism is designed tobe simple and can serve as an alternative to polycentric and single axisknee mechanisms. The knee mechanism is designed to utilize the user'sdynamics to function, which makes the knee ideal to be used bytransfemoral amputees in the K3 and K4 level. The amputees in the K3 andK4 level are more mobile and have more residual limb muscles, whichmeans they require a prosthesis that can enable them to use their motioneffectively. This knee mechanism can also be prescribed to people whohave undergone knee disarticulation [4]. Because the target populationhas more abilities, the research study tested the knee mechanism onable-bodied subjects using the prosthetic simulator depicted in FIG. 26.However, other types of users are contemplated herein as well to use thecurrent prosthesis.

In certain embodiments, the current invention may include the kneemechanism by itself or an above-knee leg prosthesis including the kneemechanism, as seen in FIG. 27 where the straight arrows indicatedirection of assembly of the individual components. The prosthesis,generally denoted by the reference numeral 10, includes femoralcomponent 12, shank component 14, knee component 16, and foot component18.

Residual limb connector 20 is coupled to femoral component 12 inoverlying relation to femoral component 12. In this particularembodiment shown in FIG. 27, residual limb connector 20 is a knee bracewith support frame in order to secure the user's knee to femoralcomponent 12. However, residual limb connector 20 can also include anadjustable prosthetic thigh, as discussed and shown previously. Further,residual limb connector 18 can be configured to fit the residual limb ofa transfemoral amputee (see FIG. 23B for example).

Shank component 14 is an elongate shaft coupled to foot component 18 inunderlying relation to knee component 16 and in overlying relation tofoot component 18. If a longer length of shank component 14 is desired,shank component 14 may include coupler 22 and/or extender 24. Thisallows shank component 14 to be extended for a greater length betweenknee component 16 and foot component 18.

Knee component 16 is coupled to femoral component 12 in underlyingrelation to femoral component 12 and is coupled to shank component inoverlying relation to shank component 14. Knee component 16 includeshousing 26 enclosing gear rack 28 and spur gear 30 with shaft andbearing assembly 32 disposed therethrough and through housing aperture34. Top side of housing 26 is typically connected to a bottom side offemoral component 12, and a bottom plate/side of spur gear 30 istypically connected to a top side of shank component 14.

Spur gear 30 translates vertically within the interior of housing 26,and similarly, shaft and bearing assembly 32 translates verticallythrough housing aperture 34. Typically, spur gear 30 and shaft andbearing assembly 32 translate vertically together. As discussedpreviously and as will be discussed in further detail as thisspecification continues, when a downward force is placed on prosthesis10 when all components thereof are disposed substantially vertically,spur gear 30 is vertically displaced in an upward direction, in turncausing spur gear 30 to mesh with gear rack 28 within housing 26. Inother words, the knee locks when the leg is straight. Conversely, whenthe downward force is released, spur gear 30 is vertically displaced ina downward position within hosing 26, in turn causing a slot or spaceddistance to form between spur gear 30 and gear rack 28 (i.e., spur gear30 and gear rack 28 are no longer meshed together). This allows shankcomponent 14 to rotate backwards relative to femoral component 12.

Femoral component 12 and shank component 14 may include stoppers 36 a,36 b to prevent shank component 14 from rotating further forward beyondthe vertical position relative to femoral component 12 (see FIG. 26). Ina vertical position, such as that seen in FIG. 26, stoppers 36 a, 36 bcontact each other at knee strike to prevent any further horizontal orvertical displacement in the undesired (i.e., unnatural) direction.Stoppers 36 a, 36 b also allow shank component 14 to assume a preciselocking position.

Foot component 18 is coupled to shank component 14 (specificallyextender 24 in FIG. 27) in underlying relation to shank component 14.Foot component 18 can take any suitable shape or configuration known inthe art, though it is preferred that foot component 18 is appropriatefor passive walking purposes. In FIG. 27, for example, foot component 18has a rollover shape to be used for such purposes. Foot component 18 isdefined by the curve followed by the foot's center of pressure pointswhen they are transformed from a general coordinate system to aknee-ankle based coordinate system [108]. In this particular embodiment,foot component 18 is a laser cut rigid piece of delrin, which is anacetal homopolymer that has a similar tensile strength as aluminum buthas a lower shear strength. This curved shape allows foot component 18to function without an ankle. The curvature of foot component 18 allowsthe user to rock forward, which simulates a downward slope and alsogives the effect of plantar-flexion, although no force is generated byfoot component 18 itself.

In certain embodiments, the knee locking mechanism is designed to be asimple passive system including only one moving part. FIGS. 28A-28D showthe various positions the knee assumes during a gait cycle. Comparingthe positions from FIGS. 28A-28D to the gait cycle depicted in FIGS.29A-29H provides a better understanding of the working of the kneemechanism. The knee assumes the position in FIG. 28A when it is locked.The locking occurs when the weight of the user is applied on themechanism, as indicated by the red arrow in the figure, which causes thespur gear of the shank to mesh with the spur gear rack of the femur. Theknee is locked while the user applies their weight on the prosthesisduring stance phase as seen in FIGS. 29H, 29G, and 29F. When the user'sweight is taken off the knee mechanism, which occurs just after toe offin FIG. 29F, the shank spur gear unmeshes with the femoral spur gearrack. The slot in the femoral housing allows the bearing of the shank totranslate ˜5 mm vertically, depicted in FIG. 28B.

This marks the beginning of the swing phase for the prosthesis when theshank is free to rotate as seen in FIG. 28C and correspondingly in FIGS.29D and 29C. The shank utilizes the motion of the user's residual limbto swing like a pendulum. When the user's residual limb reaches theextended position, the shank returns and the stopper of the shank makescontact with the stopper of the femoral housing as seen in FIG. 29B atknee strike and in closer view in FIG. 28D. Knee strike occurs justbefore heel strike and the shank assumes its position to traverse backup the slot to mesh with the femoral spur gear rack when the userapplies their weight on the prosthesis at heel strike as depicted inFIG. 29A. The knee is then back to the locked position as seen in FIG.28A. The knee strike and heel strike occur at a close interval and,hence, there is no bounce back in the prosthetic knee mechanism. Thisfacilitates the passive mechanism of the device as well. This cyclecontinues for every stride of the gait cycle.

Results and Discussion

Testing on the knee mechanism was conducted on a single subject who isexperienced with walking on the prosthetic simulator. The knee mechanismis designed to have constant periodic kinematics during every stride.The study was conducted in the CAREN system (MOTEK MEDICAL), including aBERTEC split belt treadmill, a MOOG motion base with six degrees offreedom (DOF), a ten-camera VICON (Edgewood, N.Y.) infrared motioncapture system, BERTEC force plates, and a panoramic display for fullvisual immersion. The subject's motion was captured using reflectivemarkers placed on specific locations on the subject's body. For thisstudy, the lower limb human body model [109] was utilized to positionthe reflective markers.

The data obtained from the motion capture was analyzed using a customMATLAB script. The kinematics, absolute angles during gait cycle, andangular velocity during the gait cycle were analyzed. FIGS. 30A, 31A,and 32A depict normal walking for the three (3) cases, while FIGS. 30B,31B, and 32B depict the cases with the prosthesis. The prosthesis wasworn on the subject's right leg for this study.

The plots for the knee kinematics showed an interesting trend. In thecase of normal walking, seen in FIG. 30A, the motion of both knees arerelatively symmetric and the pattern of the motion is nearly identical,which is to be expected. An expected difference in the pattern of kneekinematics was observed when the prosthesis was worn. In FIG. 30B, it isseen that the prosthetic knee is lower than the intact knee, which wasdone to accommodate the knee brace (reference numeral 20 in FIGS. 26-27)that is worn by the user to secure his knee into the prosthesis. Themotion of the prosthetic knee is seen to follow the trend that wasdescribed previously. The shank translates vertically in the femoralhousing which is seen as the space between the stance (bottom half ofthe curve) and swing (top half of the curve) phases. This regular cyclicpattern is not observed in the intact knee curve since the design is notdesigned to mimic the human knee exactly.

The differences are also seen in the absolute knee angles, measured withrespect to the ground. The shank of the knee mechanism is designed tooperate with the least amount of friction as possible. Hence, the shankis allowed to swing freely like a pendulum during swing phase. Sincethis is a passive system, the user relies on his/her dynamics and timingsteps correctly in order to perform a stable gait. This freedom tofreely rotate has produced the results as seen in FIG. 31B where it canbe seen that the prosthesis generates a greater angle during flexion, atabout 20% of the gait cycle, than the baseline right leg seen in FIG.31A. The knee angles are fairly consistent after heel strike, at about50%, in both cases which means the prosthetic knee locks successfully.It can also be seen that the intact (left) knee is compensating bygradually flexing with a quick extension.

Similar to the knee angles, the angular velocity profiles for theprosthesis show a larger magnitude of angular velocity during flexion,at about 20% of the gait cycle, and during extension, at about 50% ofthe gait cycle, as seen in FIGS. 32A-32B. The higher magnitude ofangular velocity of the prosthesis can be attributed to low resistanceto rotation of the shank. The intact knee in the prosthetic trialshowcases a more prolonged stance phase (FIG. 32B) as opposed to theprofile generated by normal walking (FIG. 32A).

CONCLUSIONS

The passive position and weight activated knee mechanism is a robustmechanism, surviving rough treatment. The results presented herein canbe seen as a preliminary step in the introduction of an alternativeprosthetic knee mechanism. This passive knee mechanism was featured in apreliminary analysis of asymmetric knee location study [110], but noanalysis was performed pertaining to the knee mechanism. A smallerversion of the knee mechanism can also be adapted to robots and bipedalwalkers, such as passive dynamic walkers, which specifically modelasymmetric gait [39, 28].

It is also contemplated herein that the current invention may includesprings and dampers in the knee mechanism to improve the dynamics andreduce forces acting on the residual limb of the amputee. Further,including traditional active systems such as motors and pistons inconjunction with springs and dampers can generate joint forces at onlyspecific stages of the gait cycle. This embodiment of semi-activesystems can retain the knee and ankle mechanisms as a low cost optionand improve battery life of the prosthesis. Adding semi-activemechanisms at the ankle and foot can aid the performance of the kneeduring push off and help in ground clearance.

The passive prosthetic knee mechanism described herein is designed toserve as an alternative to current knee systems for K3 and K4 amputees,for example. This system is designed to be customized as per therequirements of the user and is flexible when it comes to adding newcomponents to improve the control, weight distribution, and locking.

REFERENCES

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All referenced publications are incorporated herein by reference intheir entirety. Furthermore, where a definition or use of a term in areference, which is incorporated by reference herein, is inconsistent orcontrary to the definition of that term provided herein, the definitionof that term provided herein applies and the definition of that term inthe reference does not apply.

Nonlimiting Illustrative Glossary of Claim Terms

Downward facing: This term is used herein to refer to the face of acomponent presented in the inferior direction of a user of thecomponent.

Meshable relationship: This term is used herein to refer to anassociation between a gear spur and a gear rack, where the twocomponents mesh together.

Passive prosthesis: This term is used herein to refer to an apparatusthat facilitates or supports an individual's gait, where the apparatusis not powered or mechanically actuated (i.e., it does not draw energyfrom a source external to the user itself).

Passive rollover shape: This term is used herein to refer to aconfiguration of a prosthetic foot component where there is a change inthe center of pressure of the foot during walking (e.g., the foot rollsforward (without additional actuation) when a pressure is applied to thefoot by the user).

Passive: This term is used herein to refer to an apparatus, or componentthereof, that does not require power or other actuation in order tofunction.

Prosthetic: This term is used herein to refer to a structural componentbeing artificial and acting as a substitute for a user's body part(specifically a leg or portion thereof here).

Residual or impaired limb: This term is used herein to refer to anindividual's appendage that is compromised, amputated, or otherwise inneed of an aid for full functioning.

Substantially downward force: This term is used herein to refer to anyforce or weight placed by the user onto the prosthetic knee and shankcomponents sufficient to permit the spur gear to vertically displace andmesh with the gear rack.

Substantially vertical position: This term is used herein to refer to apositioning of the prosthetic shank component being sufficiently uprightto the extent that the spur gear can be vertically displaced to meshwith the gear rack when the user exerts a downward force on theprosthetic knee and shank components.

Substantially vertically-oriented: This term is used herein to refer toa positioning of the prosthetic femoral component remaining relativelyperpendicular to the ground on which the user is walking, as ittypically would be affixed to the user's impaired or residual limb. Assuch, when the user is walking, the prosthetic femoral componenttypically would only deviate a bit from its perpendicular position aswould be necessary during the user's gait.

Upward facing: This term is used herein to refer to the face of acomponent presented in the superior direction of a user of thecomponent.

Walking motion: This term is used herein to refer to the gait of anindividual, including the typical movements and rotations of theindividual's legs when walking in a forward direction.

The advantages set forth above, and those made apparent from theforegoing description, are efficiently attained. Since certain changesmay be made in the above construction without departing from the scopeof the invention, it is intended that all matters contained in theforegoing description or shown in the accompanying drawings shall beinterpreted as illustrative and not in a limiting sense.

It is also to be understood that the following claims are intended tocover all of the generic and specific features of the invention hereindescribed, and all statements of the scope of the invention that, as amatter of language, might be said to fall therebetween.

What is claimed is:
 1. A prosthetic device, comprising: a prostheticfemoral component having a first end and a second end, wherein the firstend is coupled to a residual limb connector of a user and the second endis coupled to a first end of a prosthetic knee component; the prostheticknee component having a first locked position and a second unlockedposition, the prosthetic knee component including: a housing coupled toa second end of the prosthetic knee component, the housing configured tohouse a spur gear affixed to a top portion of a prosthetic shankcomponent, and a gear rack affixed to an interior surface of thehousing, wherein the spur gear and the gear rack have a meshablerelationship when in contact with each other in the first lockedposition; the housing further including a first vertical slot and asecond vertical slot, each of the first and the second vertical slotsdisposed within the housing and include an upper limit and a lowerlimit; a shaft and bearing assembly in communication with the spur gearfacilitates rotational motion of the spur gear relative to theprosthetic shank component, wherein the shaft and bearing assembly andthe spur gear translate vertically between the upper limit and the lowerlimit of each of the first and the second vertical slots; and aprosthetic foot component having a first end and a second end, whereinthe first end of the prosthetic foot component is coupled to a bottomportion of the prosthetic shank component.
 2. The prosthetic device ofclaim 1, wherein the prosthetic femoral component comprises a firststopper and the prosthetic shank component comprises a second stopper,wherein the second stopper limits the rotation of the shank when thesecond stopper contacts the first stopper on the prosthetic femoralcomponent.
 3. The prosthetic device of claim 2, wherein limiting therotation of the shank prevents the shank from rotating further forwardthan a substantially vertical position.
 4. The prosthetic device ofclaim 1, wherein when a force is exerted from the first end of theprosthetic knee toward the second end of the prosthetic knee the forceresults in the shaft and bearing assembly moving to the upper limit ofthe vertical slots, such that when the prosthetic knee component is inthe first locked position the spur gear and the gear rack rotate withrespect to each other.
 5. The prosthetic device of claim 1, wherein whenin the second unlocked position, a force of gravity causes the shaft andbearing assembly to transition to the lower limit of the vertical slots,resulting in a space between the spur gear and the gear rack, such thatwhen the shaft and bearing assembly is at the lower limit the spur gearand the gear rack rotate independently of each other.
 6. The prostheticdevice of claim 1, wherein when the prosthetic knee component is in thesecond unlocked position, the prosthetic shank component is capable ofrotating relative to the prosthetic femoral component.
 7. The prostheticdevice of claim 1, wherein the spur gear includes a first side and asecond side, wherein meshable teeth are disposed on the first side andthe second side of the spur gear is planar.
 8. The prosthetic device ofclaim 1, wherein the prosthetic femoral component has an adjustablelength.
 9. The prosthetic device of claim 8, wherein the prostheticfemoral component comprises at least a first shaft telescopicallyreceived within a bore of a second shaft to provide for the adjustablelength.
 10. The prosthetic device of claim 1, wherein the prostheticshank component has an adjustable length.
 11. The prosthetic device ofclaim 10, wherein the prosthetic shank component comprises at least afirst shaft telescopically received within a bore of a second shaft toprovide for the adjustable length.
 12. The prosthetic device of claim10, wherein the prosthetic shank component further comprises a extensionshall to extend the length of the prosthetic shank component.
 13. Theprosthetic device of claim 1, wherein a vertical translation of theshaft and bearing assembly within the first and the second verticalslots results in less than 20 mm displacement from the upper limit andthe lower limit.
 14. The prosthetic device of claim 1, wherein thesecond end of the prosthetic foot component comprises a rollover shape.15. The prosthetic device of claim 14, wherein the rollover shapecomprises a radius of curvature of the second end of the prosthetic footcomponent decreasing toward a front side of the prosthetic footcomponent.